This invention relates to medical apparatus and methods in general, and more particularly to apparatus and methods for tissue engineering.
Tissue engineering is a truly multidisciplinary field, which applies the principles of engineering, life science, and basic science to the development of viable substitutes that restore, maintain, or improve the function of human tissues. Large-scale culturing of human or animal cells (including but not limited to skin, muscle, cartilage, bone, marrow, endothelial and stem cells) may provide substitutes to replace damaged components in humans. Naturally derived or synthetic materials are fashioned into xe2x80x9cscaffoldsxe2x80x9d that, when implanted in the body as temporary structures, provide a template that allows the body""s own cells to grow and form new tissues while the scaffold is gradually absorbed. Conventional two-dimensional scaffolds are satisfactory for multiplying cells, but are less satisfactory when it comes to generating functional tissues. For that reason, a three-dimensional (3D) bioresorbable scaffold system is preferred for the generation and maintenance of highly differentiated, tissues. Ideally, the scaffold should have the following characteristics: (i) be highly porous with an interconnected pore network for cell growth and flow transport of nutrients and metabolic waste; (ii) be biocompatible and bioresorbable, with controllable degradation and resorption rates so as to substantially match tissue replacement; (iii) have suitable surface chemistry for cell attachment, proliferation and differentiation; and (iv) have mechanical properties to match those of the tissues at the site of implantation. In vivo, the scaffold structure should protect the inside of the pore network proliferating cells and their extracellular matrix from being mechanically overloaded for a sufficient period of time. This is particularly important for load-bearing tissues such as bone and cartilage.
It is estimated that the number of bone repair procedures performed in the United States alone is over 800,000 per year. Today, skeletal reconstruction has become an increasingly common and important procedure for the orthopedic surgeon. The traditional biological methods of bone-defect management include autografting and allografting cancellous bone, applying vascularized grafts of the fibula and iliac crest, and using other bone transport techniques. Today, bone grafting is increasing and the failure rate is unacceptably high. In patients who receive various bone grafts, a failure rate ranging from 16% to 50% is reported. The failure rate of autografts is at the lower end of this range, but the need for a second (i.e., donor) site of surgery, limited supply, inadequate size and shape, and the morbidity associated with the donor site are all major concerns. Furthermore, the new bone volume maintenance can be problematic due to unpredictable bone resorption. In large defects, the body can resorb the grafts before osteogenesis is complete. Furthermore, not only is the operating time required for harvesting autografts expensive, but often the donor tissue is scarce, and there can be significant donor site morbidity associated with infection, pain, and hematoma. Allografting introduces the risk of disease and/or infection; it may cause a lessening or complete loss of the bone inductive factors. Vascularized grafts require a major microsurgical operative procedure requiring a sophisticated infrastructure. Distraction osteogenesis techniques are often laborious and lengthy processes that are reserved for the most motivated patients.
Engineering osseous tissue by using cells in combination with a synthetic extracellular matrix is a new approach compared to the transplantation of harvested tissues. Numerous tissue-engineering concepts have been proposed to address the need for new bone graft substitutes. One potentially successful repair solution seeks to mimic the success of autografts by removing cells from the patient by biopsy and then growing sufficient quantities of mineralized tissue in vitro on implantable, 3D scaffolds for use as functionally equivalent autogenous bone tissue. In this way, reproducing the intrinsic properties of autogenous bone material creates an ideal bony regeneration environment, which includes the following characteristics: (i) a highly porous, 3D architecture allowing osteoblast, osteoprogenitor cell migration and graft revascularization; (ii) the ability to be incorporated into the surrounding host bone and to continue the normal bone remodeling processes; and (iii) the delivery of bone-forming cells and osteogenic growth factors to accelerate healing and differentiation of local osteoprogenitor cells.
Research on cartilage goes back more than 250 years when Hunter stated in 1743, xe2x80x9cFrom Hippocrates to the present age, it is universally known that ulcerated cartilage is a troublesome thing and that once destroyed, is not repairedxe2x80x9d. Since then, a substantial amount of research has been conducted on hyaline cartilage, fibrocartilage and elastic cartilage, with significant advances in our understanding of the development biology and biological cartilage repair process being made over the past four decades. However, the cartilage repair and regeneration response is limited in terms of form and function. While many surgical techniques and drug treatments have been proposed over the past 10 years, none have successfully regenerated long-lasting cartilage tissue to replace damaged or diseased cartilage from a clinical point of view. In fact, most of the surgical interventions to repair damaged cartilage have been directed toward the treatment of clinical symptoms rather than the regeneration of hyaline cartilage, such as pain relief and functional restoration of joint structures and articulating surfaces.
Since cartilage is limited in its ability to repair, significant efforts have also been dedicated to growing cartilage ex vivo or to supplement implants with cells to improve healing. Modern tissue engineering approaches, such as the transplantation of isolated and seeded chondrocytes in combination with bioresorbable polymeric scaffolds of synthetic and natural origin, have recently demonstrated significant clinical potential for the regeneration of different cartilaginous tissues. The success of chondrocyte transplantation and/or the quality of neocartilage formation strongly depends on the specific cellcarrier material.
Current research has largely focused on chondrocyte interaction with biodegradable polymers and devices that are FDA approved, namely, foams and textiles made of poly (glycolic acid) (PGA), poly (L-lactic acid). (PLLA) and their copolymer, poly (DL-lactic-co- glycolic acid) (PLGA). The physical and mechanical scaffold properties can have a profound effect on the healing response of the cartilage. Furthermore, the proper mechanical environment of chondrocytes and their matrix is essential to obtain a structurally and biochemically functional tissue. PGA and PLA and their copolymers have frequently been chosen for tissue engineering applications because their degradation can be tailored. In a highly porous configuration, however, their mechanical properties may be limited.
The clinical goals for craniofacial skeletal reconstruction are multifaceted. Aesthetic and functional considerations often dictate the use of malleable implant materials. However, in most cases these three-dimensional shaped implants must also provide immediate structural integrity. The host-graft interface should not produce an immunological or inflammatory response to minimize peri-implant morbidity.
Successful craniofacial surgical experience with the patient""s own bone has made it the graft material against which all others are measured. Unfortunately, autografted bone is limited in amount and desired morphology. In addition, the use of the patient""s own bone is associated with donor-site morbidity and graft resorption. When autogenous tissue is not available, or its use is limited because of defect size or shape, a variety of alloplastic materials are used for craniofacial reconstruction. Metallic, ceramic, and synthetic polymer materials are readily available. However, all metals, most ceramics, and many polymers are not designed to degrade and resorb, and the potential for replacement by host tissue does not exist. Furthermore, if the soft tissues surrounding an alloplastic material implant site were previously more firmly attached to an underlying bony surface, they may now move relative to the implanted alloplastic material. These undesirable mechanical interactions can result in seroma formation and soft tissue inflammation. Clinically, foreign body reactions are observed and long-term problems include a minimal potential for new bone growth and poor remodeling adjacent to the reconstructed area.
Stress shielding of nearby bone from the alloplastic transplant, due to the mismatch of the mechanical properties of the synthetic implant and host bone, can lead to local tissue atrophy and necrosis. Currently, polymethylmethacrylate (PMMA) is probably the most commonly used alloplastic material for reconstruction of significant craniofacial, defects. Often referred to as bone cement, PMMA is used on a routine basis as an alternative to bone autografts in reconstructing cranial defects because: (1) it can withstand the mechanical forces of the craniofacial skeleton, thus providing support and protection to underlying soft tissue structures; (2) it is an inexpensive implant material when compared to an autograft; and (3) it demonstrates sufficient biocompatibility with adjacent soft tissues.
Although primary tissue reaction of PMMA has been rarely observed, secondary foreign body reactions have been reported. Hence, the threshold for indirect infection at the surgical site is lowered. In general, PMMA is used to fill bony defects by one of two methods: it can be polymerized in situ or it can be molded and placed into the defect. Due to in vivo polymerization of PMMA, a significant exothermic reaction takes place with temperatures reaching 81xc2x0 C. The reaction can result in thermal injury to, and necrosis of, adjacent host bone, dura, and other soft tissues.
Many different processing techniques have been developed to design and fabricate three-dimensional (3D) scaffolds for tissue engineered implants. These conventional techniques include fiber bonding, solvent casting, particulate leaching, membrane lamination, melt molding, temperature-induced phase separation (TIPS), and gas foaming. A wide range of scaffold characteristics, such as porosity and pore size, has been reported using such fabrication techniques. However, there are numerous drawbacks to using such scaffolds for tissue engineering applications. For one thing, the pores are not fully interconnected due to the formation of skin-layers during solvent evaporation. And the pore size varies, as it is difficult to ensure that the porogens are well-dispersed and not agglomerated to form bigger particles. In addition, the thickness and length of the pore walls and edges vary, depending on the solvent evaporation rate. Also, the scaffolds cannot be made with thick sections inasmuch as deeply embedded porogens become too distant from the surface and residual porogens may be left in the final structure. And the use of organic solvents requires careful and complete removal of residual solvents prior to clinical usage. In addition, the aforementioned conventional scaffold fabrication techniques do not allow the fabrication of a 3-D scaffold with a varying multiple layer design. Such matrix architecture is advantageous in instances where be tissue engineers want to grow a bi- or multiple tissue interface, e.g., an articular cartilage/bone transplant. Rapid Prototyping machines (RP), such as Fused Deposition Modelling (FDM) and 3D Printing (3DP) which build a physical model by depositing layers of a material one at a time, allow such a design.
Rapid Prototyping (RP) is a technology that produces models and prototype parts from 3D computer-aided design (CAD) model data and model data created from 3D object digitizing systems. Unlike milling machines, which are subtractive in nature, RP systems join together liquid, powder and sheet materials to form parts. Layer by layer, REP machines fabricate plastic, wood, ceramic and metal objects using thin horizontal cross sections directly from a computer generated model. Rapid prototyping technologies allow the development of manufacturing approaches to create porous scaffolds that mimic the microstructure of living materials.
The application of rapid prototyping (RP) technologies in medicine has, until recently, been largely restricted to the surgical planning and simulation for reconstructive surgery, e.g., to the fabrication of prosthesis models such as cranial titanium plates to repair skull defects.
3DP has been used to process bioresorbable scaffolds for tissue engineering applications. The technology is based on the printing of a binder through a print head nozzle onto a powder bed. However, the removal of entrapped powder is typically quite difficult. The entire 3DP process is performed under room temperature conditions whereas FDM uses a thermoplastic polymer. Hence 3DP allows the incorporation of biological agents, such as cells, growth factors, and so forth, without inactivation if non-toxic binders such as water can be used. Unfortunately, aliphatic polyesters can generally only be dissolved in highly toxic solvents such as chloroform and methylene chloride. To date, only bioresorbable scaffolds without biological agents within the polymer matrix and in combination with particle leaching have been processed by 3DP. In addition, the mechanical properties and accuracy of the specimens manufactured by 3DP still have to be significantly improved.
Other RP technologies, such as stereolithography (SLA) and Selective Laser Sintering (SLS), pose significant material (non-bioresorbable) constraints for the manufacture of tissue engineering scaffolds. Ballistic Particular Manufacturing (BPM) is also limited by the lack of suitable biodegradable materials. Metal and metal composites have been proposed as processing materials for BPM.
Fused Deposition Modelling (FDM) is an additive manufacturing process that forms 3D objects through the extrusion and deposition of individual layers of thermoplastic materials. It begins with the creation of a conceptual CAD model on the computer. The model is imported into software (e.g., the QuickSlice(trademark) software offered by Stratasys Inc. of Eden Prairie, Minn.) which mathematically slices the conceptual model into horizontal layers. This is followed by the creation of deposition paths within each sliced layer. The tool path data is then downloaded to the FDM machine for scaffold fabrication. A software package (e.g., SupportWorks offered by Straftasys Inc.) automatically generates supports if needed. The FDM system operates in the X, Y and Z axes. In effect, it draws the designed model one layer at a time. The FDM method involves the melt extrusion of filament materials through a heated nozzle and deposition as thin solid layers on a platform. The nozzle is positioned on the surface of a build platform at the start of fabrication. It is part of the extruder head (FDM head), which also encloses a liquefier to melt the filament material fed through two counter-rotating rollers. Each layer is made of xe2x80x9craster roadsxe2x80x9ddeposited in the x and y directions. A xe2x80x9cfill gapxe2x80x9d can be programmed between the roads to provide horizontal channels. Subsequent layers are deposited with the x-y direction of depositionxe2x80x94the xe2x80x9craster anglexe2x80x9d programmed to provide different lay-down patterns.
Thermoplastic polymer material feeds into the temperature-controlled FDM extrusion head, where it is heated to a semi-liquid state. The head extrudes and deposits the material in ultra-thin layers onto a fixtureless base. The head directs the material into place with precision. The material solidifies, laminating to the preceding layer. Parts are fabricated in layers, where each layer is built by extruding a small bead of material, or road, in a particular lay-down pattern, such that the layer is covered with the adjacent roads. After a layer is completed, the height of the extrusion head is increased and the subsequent layers are built to construct the part. Usually, FDM is used to fabricate solid models. For the purpose of fabricating porous structures, a positive value is applied to the raster fill gap to impart a channel within a build layer. Arranged in a regular manner, the channels are interconnected even in three dimensions. The layer by layer fabrication allows design of a pore morphology which varies across the scaffold structure. At present, only a few non-resorbable polymeric materials, such as polyamide, ABS, resins, etc. are used in the FDM RP systems.
Poly(caprolactone) (PCL) is a semicrystalline, bioresorbable polymer belonging to the aliphatic polyester family. It has favorable properties for thermoplastic processing. It has a low glass transition temperature (Tq) of xe2x88x9260xc2x0 C., a melting point (Tm) of 60xc2x0 C. and a high decomposition temperature of 350xc2x0 C., with a wide range of temperatures which allow extrusion. At present, PCL is regarded as a soft and hard tissue-compatible bioresorbable material.
The present invention uses FDM to process a bioresorbable polymer, polycaprolactone (PCL), as well as a bioresorbable composite of two biomaterials, synthetic polymer (PCL) and ceramic, to meet all the criteria for use in tissue engineering applications.
The present invention relates to the use of FDM to construct three-dimensional (3D) bioresorbable scaffolds from polycaprolactone (PCL), and from composites of PCL and ceramics, such as tricalcium phosphate (TCP) and hydroxyapatite (HA), with specific lay-down patterns that confer the requisite properties for tissue engineering applications.
The 3D polymer matrix has degradation and resorption kinetics of 6 to 12 months and,the capability to maintain a given space under biomechanical stress/loading for 6 months. Incorporation of a bioresorbable ceramic in the bioresorbable, synthetic and natural polymer produces a hybrid/composite material support triggering the desired degradation and resorption kinetics. Such a composite material improves the biocbmpatibility and hard tissue integration: the HA/TCP particles, which are embedded into the synthetic polymer matrix, allow for increased initial flash spread of serum proteins compared to the more hydrophobic polymer surface. Furthermore, the basic resorption products of the HA/TCP help buffer the acidic resorption by-products of the aliphatic polyester and thereby help to avoid the formation of an unfavorable environment for the hard tissue cells due to a decreased pH.
The resulting scaffolds have applications in tissue engineering such as tissue engineering bone and cartilage.
In one form of the invention, there is provided a method for fabricating a filament for use in tissue engineering, the method comprising:
providing a polycaprolactone material;
melting the polycaprolactone material at a first given temperature to form a polycaprolactone melt;
holding the temperature of the polycaprolactone melt at the first given temperature for a given amount of time;
lowering the temperature of the polycaprolactone melt from the first given temperature to a second given temperature after the step of holding the temperature of the polycaprolactone melt at the first given temperature for the given amount of time;
extruding the polycaprolactone melt through a fiber-spinning machine, the fiber-spinning machine having spinnerets with a die exit of a given diameter, a piston set at a given speed, and a vertical drop of a given distance from the die exit to a cooling material positioned below the die exit, wherein the combination of the second given temperature, the given die exit diameter, the given piston speed, and the given distance of the vertical drop produces the filament with a given diameter for use in tissue engineering.
In another form of the invention, there is provided a method for fabricating a filament for use in tissue engineering, the method of fabricating the filament comprising:
providing polycaprolactone pellets;
melting the polycaprolactone pellets at about 190xc2x0 C. to form a polycaprolactone melt;
holding the temperature of the polycaprolactone melt at about 190xc2x0 C. for about 15 minutes;
lowering the temperature of the polycaprolactone melt from the first given temperature to about 140xc2x0 C. after the step of holding the temperature of the polycaprolactone melt at about 190xc2x0 C. for about 15 minutes; and
extruding the polycaprolactone melt through a fiber-spinning machine, the fiber-spinning machine having spinnerets with a die exit of about 1.63 mm, a piston set at about 10 mm/min, and a vertical drop of about 40 mm from the die exit to water positioned below the die exit, wherein the combination of the lower temperature of about 140xc2x0 C., the die exit diameter of about 1.63, the piston speed of about 10 mm/min, and the vertical drop of about 40 mm produces the filament with a given diameter for use in tissue engineering;
wherein the given diameter of the filament corresponds to drive wheels of an unmodified Fused Deposition Modeling (FDM) system;
wherein the filament is configured to have a constant diameter; and
wherein the filament is vacuum-dried and kept in a dessicator prior to usage.
And in another form of the invention, there is provided a method for fabricating a filament for use in tissue engineering, the method comprising:
providing a polycaprolactone material drying the polycaprolactone material at a first given temperature for a first given amount of time to form a dried polycaprolactone material;
combining the dried polycaprolactone material with a HA and methylene chloride mixture to form a PCL/HA blend;
stirring the PCL/HA blend at a second given temperature for a second given amount of time to form a solvent mixture;
casting the solvent mixture on a tray at a third given temperature for a third given amount of time to evaporate the solvent mixture to form a PCL/HA composite foam material;
melting the PCL/HA composite foam material at a fourth given temperature to form a PCL/HA melt;
holding the temperature of the PCL/HA melt at the fourth given temperature for a fourth given amount of time;
lowering the temperature of the PCL/HA melt from the fourth given temperature to a fifth given temperature after the step of holding the temperature of the PCL/HA melt at the fourth given temperature for the fourth given amount of time; and
extruding the PCL/HA melt through a fiber-spinning machine, the fiber spinning machine having spinnerets with a die exit of a given diameter, a piston set at a given speed, and a vertical drop of a given distance from the die exit to a cooling material positioned below the die exit, wherein the combination of the fifth given temperature, the given die exit diameter, the given piston speed, and the given distance of the vertical drop produces the filament with a given diameter for use in tissue engineering.
In another form of the invention, there is provided apparatus for use in tissue engineering, the apparatus comprising:
a scaffold structure being formed of a plurality of horizontal layers of material;
vertical walls forming each of the plurality of horizontal layers of material, the walls of each layer of the plurality of horizontal layers each having a height, each being horizontally separated from one another, and defining an orientation;
adjacent pairs of the vertical walls of each of the plurality of horizontal layers of material forming channels therebetween, the channels having a depth and a width created by the height of the walls and the horizontal separation of the adjacent pairs of the vertical walls, respectively;
adjacent layers in the plurality of horizontal layers of material being in different orientations to one another wherein the orientation defined by adjacent ones of the each layer of the walls of the plurality of horizontal layers differ from one another, the different orientations providing a group of cross-points to allow adhesion between the adjacent layers and providing interconectivity between the channels throughout the scaffold structure.
And in another form of the invention, there is provided a method for fabricating a customized scaffold structure for use in tissue engineering for an individual patient, the method comprising:
obtaining a digital scan of an anatomical component of the individual patent;
obtaining a desired zone of the digital scan of the anatomical component of the individual patent;
converting the desired zone of the digital scan of the anatomical component of the individual patent to an Fused Deposition Modeling (FDM) system compatible format;
slicing the desired zone of the digital scan of the FDM system compatible format into multiple layers so as to create a sliced model of the customized scaffold structure for fused deposition modeling;
creating tool path data for fused deposition modeling using the sliced model of the customized scaffold structure;
exporting the tool path data of the sliced model of the customized scaffold structure to a Fused Deposition Modeling (FDM) system; and
creating the customized scaffold structure using the tool path data of the sliced model of the customized scaffold structure and the Fused Deposition Modeling (FDM) system.